Nuclear imaging is a unique specialty wherein radiation is used to acquire images that show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the subject, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions that emanate from the body. One or more detectors are used to detect the emitted gamma photons, and the information collected from the detector(s) is processed to calculate the position of origin of the emitted photon from the source (i.e., the body organ or tissue under study). The accumulation of a large number of emitted gamma positions allows an image of the organ or tissue under study to be displayed.
In conventional nuclear imaging arrangements, collimators are used in a wide variety of equipment in which it is desired to permit only beams of radiation emanating along a particular path to pass a selected point or plane. Collimators are frequently used in nuclear imagers to ensure that only radiation beams passing along a direct path from the known radiation source strike the detector thereby minimizing detection of beams of scattered or secondary radiation.
Particularly in nuclear imagers used for preclinical analysis or for non-destructive evaluation procedures, it is important that only radiation emanating from a known source and passing along a direct path from that source be detected and processed by the imaging equipment. If the detector is struck by undesired radiation such as that passing along non-direct paths to the detector, performance of the imaging system can be degraded.
Two principal types of collimators have been used in nuclear imaging. The predominant type of collimation is the parallel-hole collimator. This type of collimator contains hundreds of parallel holes drilled or etched into a very dense material such as lead. The parallel-hole collimator accepts only photons traveling perpendicular to the scintillator surface, and produces a planar image of the same size as the source object. In general, the resolution of the parallel-hole collimator increases as the holes are made smaller in diameter and longer in length.
The conventional pinhole collimator typically is cone-shaped and has a single small hole drilled in the center of the collimator material. The pinhole collimator generates a magnified image of an object in accordance with its acceptance angle, and is primarily used in studying small organs such as the thyroid or localized objects such as a joint. The pinhole collimator must be placed at a very small distance from the object being imaged in order to achieve acceptable image quality. Pinhole collimators offer the benefit of high magnification of a single object, but lose resolution and sensitivity as the field of view (FOV) gets wider and the object is farther away from the pinhole.
U.S. Pat. No. 7,166,846, assigned to the same assignee herein and incorporated herein by reference in its entirety, discloses a multi-pinhole collimator nuclear medical imaging detector that divides a target object space into many non-overlapping areas and projects a minified image of each area onto a segmented detector, where each segment functions as an independent detector or imaging cell.
Other known types of collimators include the slant-hole collimator, converging and diverging collimators, and the fan beam collimator. The slant-hole collimator is a variation of the parallel-hole collimator but with all holes slanted at a specific angle. This type of collimator is positioned close to the body and produces an oblique view for better visualization of an organ whose line of sight may be partially blocked by other parts of the body. The converging collimator has holes that are not parallel but instead are focused toward the organ, with the focal point being located in the center of the field of view. The image appears larger at the face of the scintillator using a converging collimator. A diverging collimator results by reversing the direction of the converging collimator. The diverging collimator is typically used to enlarge the FOV, such as would be necessary with a portable camera having a small scintillator. The fan beam collimator is typically used with a rectangular camera head to image smaller organs. The holes are parallel when viewed from one direction and converge when viewed from another direction. The fan beam collimator allows the maximum surface of the crystal to be used to capture imaging data. In most applications, the choice of collimation represents a trade-off between the size of the FOV and the sensitivity and spatial resolution required to properly visualize the target object or organ.
Collimators are positioned to substantially absorb the undesired radiation before it reaches the detector. The collimator includes (or is manufactured from) a relatively high atomic number material and the collimator is positioned so that undesired radiation strikes the body of the collimator and is absorbed before being able to strike the detector. In a typical detector system the collimator includes barriers associated with the detector and located in the direction of the radiation source. An example exists in radiation imaging systems used for medical diagnosis which use a small point source of radiation to expose the patient under examination. The radiation passes through the patient and strikes a radiation detector that is oppositely positioned.
Conventional single photon imaging systems with parallel-hole collimation use large area (on the order of 2000 cm2) monolithic scintillation detectors, and typically have an intrinsic spatial resolution of approximately 3.5 mm FWHM (Full Width Half Maximum). Such detectors are made either of sodium iodide crystals doped with thallium (NaI(Tl)), or cesium iodide (CsI). Scintillations within the NaI crystal caused by absorption of a gamma photon within the crystal, result in the emission of a number of light photons from the crystal. The scintillations are detected by an array of photomultiplier tubes (PMTs) in close optical coupling to the crystal surface.
The intrinsic spatial resolution is primarily determined by the size of the PMTs. The design of the parallel-hole collimator (i.e., the length and diameter of the collimator holes) fixes the system resolution, and represents a trade-off between sensitivity (i.e., the number of detected gamma rays) and spatial resolution (i.e., sharpness of the image) of the imaged target object. The system spatial resolution is a quadrature sum of the geometric resolution of the collimator and the intrinsic resolution of the camera. In most clinical imaging studies, the predominant spatial resolution achieved is determined by the geometric resolution of the collimator, and thus there has not been a strong incentive to increase the intrinsic spatial resolution of the gamma camera.
Conventional commercial gamma cameras are designed to minimize cost by using the largest possible size PMTs, and thus achieve an intrinsic spatial resolution of about 3.5 mm FWHM. However, recent detector technology has enabled the design of small gamma cameras with intrinsic spatial resolution of less than 1 mm FWHM. Thus, there exists a need in the art for improvements in collimator technology to take advantage of such increased intrinsic spatial resolution in the development of new commercial gamma cameras.
In the instance of SPECT scanning, a subject (patient) is infused with a radioactive substance that emits radioactive or gamma rays. Conventionally, a gamma camera includes a transducer to receive the gamma rays and record an image therefrom. In order for the image to be a true representation of the subject being investigated, a collimator having collimating apertures is positioned between the transducer and the subject to screen out all of the radioactive rays except those directed along a straight line through the collimating apertures between a particular part of the subject and a corresponding particular part of the transducer. Traditionally, the collimator is made of a radiation opaque material such as lead, and collimating apertures have been formed therein by various means such as drilling holes therethrough.
In conventional SPECT system designs, lead collimator gamma cameras have been supported on gantries that rotate the camera head through an angular range of one hundred eighty or three hundred sixty degrees around the patient. One drawback associated with this requirement however, is that such gantry systems are relatively expensive subsystems of the diagnostic tool because they must be capable of providing rapid rotation of the large and heavy camera heads through very precise orbits about the patient. Further, rotating gantries require a large degree of space for the actual unit as well as for full operational ability. This is especially problematic with preclinical SPECT wherein lab space limitations are more prevalent. As a result, one object of the present invention is to accommodate the use of space and money-saving simplified gantries, without sacrificing image quality.